The ability to produce images of the inside of a living organism without invasive surgery has been a major advancement in medicine over the last one hundred years. Imaging techniques such as X-ray computed tomography (CT) and magnetic resonance imaging (MRI) have given doctors and scientists the ability to view high-resolution images of anatomical structures inside the body. While this has led to advancements in disease diagnosis and treatment, a large set of diseases causes changes in anatomical structure only in the late stages of the disease or never at all. This has given rise to a branch of medical imaging that captures certain metabolic activities inside a living body. Positron emission tomography (PET) is in this class of medical imaging.
Positron Emission Tomography
PET is a medical imaging modality that takes advantage of radioactive decay to measure certain metabolic activities inside living organisms. PET imaging systems comprise three main components, indicated schematically in FIG. 1, a radioactive tracer that is administered to the subject to be scanned, a scanner that is operable to detect the location of radioactive tracer (indirectly as discussed below), and a tomographic imaging processing system.
The first step is to produce and administer a radioactive tracer 90, comprising a radioactive isotope and a metabolically active molecule. The tracer 90 is injected into the body to be scanned 91. After allowing time for the tracer 90 to concentrate in certain tissues, the body 91 is positioned inside the scanner 92. The radioactive decay event for tracers used in conventional PET studies is positron emission. Radioactive decay in the tracer 90 emits a positron e+. The positron e+ interacts with an electron e− in the body in an annihilation event that produces two 511 KeV anti-parallel photons or gamma photons γ. The scanner 92 detects at least some of the 511 KeV photons γ generated in the annihilation event.
The scanner 92 includes a ring of sensors and front-end electronics that process the signals generated by the sensors. The sensors typically comprise scintillator crystals or scintillators 93 and photodetectors 94 (e.g., photomultiplier tubes (PMT), silicon photomultipliers (SiMP) or avalanche photo diodes (APD)). The scintillator crystal 93 interacts with the 511 KeV gamma photons γ to produce many lower-energy photons, typically visible light photons. The photodetector 94 detects the visible light photons and generate a corresponding electrical pulse or signal. The electric pulses are processed by front-end electronics to determine the parameters or characteristics of the pulse (i.e., energy, timing). Unless the context implies otherwise, for convenience references to a photodetector herein will be understood to include any mechanism or device for detecting gamma photons such as 511 KeV photons and producing lower-energy photons such as visible light photons in response.
Finally, the data is sent to a host computer 95 that performs tomographic image reconstruction to turn the data into a 3-D image.
Radiopharmaceutical
To synthesize the tracer 90, a short-lived radioactive isotope is attached to a metabolically active molecule. The short half-life reduces the subject's exposure to ionizing radiation, but generally requires the tracer 90 be produced close to the scanner. The most commonly used tracer is fluorine-18 flourodeoxyglucose ([F-18]FDG), an analog of glucose that has a half-life of 110 minutes. [F-18]FDG is similar enough to glucose that it is phosphorylated by cells that utilize glucose, but does not undergo glycolysis. Thus, the radioactive portion of the molecule becomes trapped in the tissue. Cells that consume a lot of glucose, such as cancers and brain cells, accumulate more [F-18]FDG over time relative to other tissues.
After sufficient time has passed for the tissue of interest to uptake enough tracer 90, the scanner 92 is used to detect the radioactive decay events, i.e., by detecting the 511 KeV photons. When a positron is emitted, it typically travels a few millimeters in tissue before it annihilates with an electron, producing two 511 KeV photons directed at 180°±0.23° from one another.
Photon Scintillation
Most of the 511 KeV photons will pass through the body tissue (and other materials) without significant interaction. While this typically allows the photon to travel through and exit the body, the gamma photons are difficult to detect. Photon detection is the task of the scintillator 93. A scintillator 93 absorbs gamma photons and emits lower energy photons, typically visible light photons. A scintillator 93 can be made from various materials including plastics, organic and inorganic crystals, and organic liquids. Each type of scintillator has a different density, index of refraction, timing characteristics, and wavelength of maximum emission.
In general, the density of the scintillator crystal determines how well the material stops the gamma photons. The index of refraction of the scintillator crystal and the wavelength of the emitted light affect how easily light can be collected from the crystal. The wavelength of the emitted light also needs to be matched with the device that will turn the light into an electrical pulse (e.g., the photodetector) in order to optimize the efficiency. The scintillator timing characteristics determine how long it takes the visible light to reach its maximum output (rise time) and how long it takes to decay (decay time). The rise and decay times are important because the longer the sum of these two times, the lower the number of events a photodetector can handle in a given period, and thus the longer the scan will take to get the same number of counts. Also, the longer the timing characteristics, the greater the likelihood that two events will overlap (pile-up) and data will be lost.
The 511 KeV photons may undergo two types of interactions within the scintillator 93: Compton scattering, wherein the photon will lose energy and change direction, and photoelectric absorption. For example, a particular gamma photon may (i) experience photoelectric absorption in its first interaction in the scintillator crystal, (ii) undergo Compton scattering one or more times within the crystal prior to photoelectric absorption, or (iii) may undergo Compton scattering one or more times within the crystal before being ejected from the crystal.
Photodetectors
Attached to the scintillator 93 are electronic photodetectors 94 that convert the visible light photons from the scintillator 93 into electronic pulses. The two most commonly used devices are PMTs and APDs. A PMT is a vacuum tube with a photocathode, several dynodes, and an anode that has high gains to allow very low levels of light to be detected. An APD is a semiconductor version of the PMT. Another technology that is currently being studied for use in PET scanners are SiPMs. SiPMs (also called Geiger-Mode APDs (GM-APD)) comprise an array of semiconducting photodiodes that operate in Geiger mode so that when a photon interacts and generates a carrier, a short pulse of current is generated.
In an exemplary SiPM, the array of photodiodes comprises about 103 diodes per mm2. All of the diodes are connected to a common silicon substrate so the output of the array is a sum of the output of all of the diodes. The output can therefore range from a minimum wherein one photodiode fires to a maximum wherein all of the photodiodes fire. This gives these devices a linear output even though they are made up of digital devices.
Image Construction
Referring to FIG. 1, a PET system acquires data as follows: photodetector 94 data is filtered with a low-pass filter 96, digitized with an analog to digital converter 97, and the digitized data is initially processed with field programmable gate arrays (FPGAs) 98.
The analog pulses generated by the photodetector 94 contain the information used to create a PET image. The analog pulses are processed to extract start time, location, and total energy. The apparatus for performing this initial processing is referred to as the front-end electronics, and includes the filters 96, ADCs 97, and FPGAs 98. The analog pulse received from the photodetector 94 is filtered with the low pass filter 96 to remove noise, and then digitized with the ADC 97, for processing by the FPGA 98. Another consideration is the number of inputs to the FPGA 98. Very fast ADCs have a parallel output which would require 10-12 bits per channel, with tens to hundreds of channels per FPGA. The number of inputs would thus outnumber the amount that even modern FPGAs can handle. Therefore, current systems use serial output ADCs 97, which limits the sampling rate to around 100 MSPS. However, for systems requiring fewer ADCs per FPGA, faster ADCs can be used to achieve better timing resolution.
After the analog pulse data is digitized, the requisite pulse parameters can be extracted in the FPGA 98. A total pulse energy, for example, may be obtained by summing the samples of the pulse values and subtracting out the baseline (the output value of the ADC 97 without an input pulse).
Finally, the data is sent to a host computer 95 that performs tomographic image reconstruction to turn the data into a 3-D image.
An important advantage of PET imaging is that the annihilation event produces two substantially anti-parallel 511 KeV photons. Therefore, with detectors disposed around the body being imaged, two detection events may be observed at roughly the same time (coincident events) in two oppositely-disposed sensors. The annihilation event producing the 511 KeV photons will be located somewhere on the line connecting the two photon detection points. The line connecting two coincident events is referred to as the line of response (LOR). When enough coincidental events have been detected, image reconstruction can begin. Essentially the detected events are separated into parallel lines of response (interpreted path of photon pair) that can be used to create a 3-D image using computer tomography. Methods for creating images using computer tomography are well known in the art. It will be appreciated that the accuracy of the 3-D PET images is dependent on the accuracy of the estimated LORs.
While PET, MRI, and CT are all common medical imaging techniques, the information obtained from the different modalities is quite different. MRI and CT give anatomical or structural information. That is, they produce a picture of the inside of the body. This is great for problems such as broken bones, torn ligaments or anything else that presents as abnormal structure. However, MRI and CT do not indicate metabolic activity. This is the domain of PET. The use of metabolically active tracers means that the images produced by PET provide functional or biochemical information.
Oncology (study of cancer) is currently the most common application of PET. Certain cancerous tissues metabolize more glucose than normal tissue. [F-18]FDG is close enough to glucose that cancerous cells readily absorb it and, therefore, they have high radioactive activity relative to background tissue during a scan. This enables a PET scan to detect some cancers before they are large enough to be seen on an MRI scan. PET scan information is also very useful for monitoring treatment progression as the quantity of tracer uptake can be tracked over the progression of the therapy. If a scan indicates lower activity in the same cancerous tissue after therapy, it indicates the therapy is working.
PET is also useful in neurology (study of the nervous system) and cardiology (study of the heart). An interesting application in neurology is the early diagnosis of Parkinson's disease. Tracers have been developed that concentrate in the cells in the brain that produce dopamine, a neurotransmitter. In patients with Parkinson's disease, neurons that produce dopamine reduce in number. Therefore, a scan of a Parkinson's patient would have less activity than a healthy patient's. This can lead to early diagnosis, since many of the other early signs of Parkinson's are similar to other diseases.
Detector Design
In a typical PET scanner 92, detectors comprising scintillators 93 and photodetectors 94 are used to perform data acquisition for image construction. FIG. 2 illustrates schematically two opposed detector modules 100A and 100B, each module comprising an array of detector elements 102. Two common types of detector elements are the continuous miniature crystal element (cMiCe) and the discrete miniature crystal element (dMiCE). The cMiCe and dMiCE detectors operate using different scintillator configurations and, in some cases, hardware and software for processing detector data.
The cMiCe detector is a continuous crystal having photodetectors mounted on at least one surface of the crystal in a specific configuration (e.g., a grid array). Because scintillation events occur in continuous crystal, and photons produced by the scintillation may spread over an area greater than that of a single photodetector, photons may be detected by multiple detectors for a single event. The light is not contained, artificially (as in a dMiCE detectors), in an area above a detector, and, so, hardware and software must be provided in the PET system to capture, analyze, and process the received photon signals from the annihilation events so as to construct an image.
The dMiCE detector modules have traditionally been used to achieve high spatial resolution for small-animal PET scanners. A dMiCE crystal pair detector element 102 is illustrated in FIG. 3, which comprises a pair of crystals 106, 106′, each crystal having a photodetector (such as a MAPD) 108, 108′ at a distal end. A triangular reflector 110 is disposed between the two crystals 106, 106′, such that the signals from the MAPDs 108, 108′ resulting from gamma photon interactions within the crystal pair will depend on the location of the interaction within the crystal pair.
Referring to FIG. 2, the detector modules 100A and 100B may represent any two detector modules that are within the relevant field of view in a PET scanner. A source of gamma photons 104 (e.g., from the object under test) between the two detector modules 100A, 100B, for example annihilation events as discussed above, may produce a coincident detection event in the modules. In the present application a “coincident” detection event refers to a detection event in two modules that are close enough in time to be considered coincident for purposes of PET. For example, detection events in two detector modules 100A, 100B that are within a relevant field of view may be considered coincident if they occur within a ten nanosecond period. The actual time span threshold will depend on the particular apparatus and application. Several factors may contribute to the gamma photons in an annihilation event having slightly differing detection times, including the particular location of the annihilation event within the scanner, finite timing resolution of the detector, etc.
An exemplary method and apparatus for data acquisition in PET systems is disclosed in copending U.S. Ser. No. 12/264,093, published on Sep. 10, 2009, in U.S. Patent Application Publication No. 2009/0224158, which is hereby incorporated by reference in its entirety.
As PET systems become more complex and of larger scale, the massive amounts of data generated by the photodetector arrays used in the systems creates both hardware and software problems during processing. There remains a great need for continued improvement in the cost, efficiency, and accuracy of PET systems, as well as the reduction in complexity and amount of data that must be handled by a software-based computer system.